Magnetic resonance imaging (MRI) is a technique using magnetic resonance phenomena for imaging. In magnetic resonance phenomena, a nucleus containing a single proton (e.g., the proton of a hydrogen nucleus prevalent in the human body) has a spinning movement similar to a small magnet. The spin axes of these small magnets do not follow a certain rule. If an external magnetic field is applied, these small magnets will be rearranged according to the magnetic force lines of the external magnetic field. For example, the small magnets may be arranged in two directions that are parallel or antiparallel to the magnetic force lines of the external magnetic field. The direction parallel to the magnetic force lines of the external magnetic field is called a positive longitudinal axis. The direction antiparallel to the magnetic force lines of the external magnetic field is called a negative longitudinal axis. The nucleus has only a longitudinal magnetization component, and the longitudinal magnetization component has both direction and amplitude. Nuclei in the external magnetic field may be excited by a radio-frequency (RF) pulse with a specific frequency to make the spin axes of the nuclei deviate from the positive longitudinal axis or the negative longitudinal axis, thereby producing resonance and giving rise to a magnetic resonance phenomenon. After the spin axes of the excited nuclei deviate from the positive longitudinal axis or the negative longitudinal axis, the nuclei have a transverse magnetization component. After stopping transmission of the radio-frequency pulse, the excited nuclei transmit echo signals to release the absorbed energy piecemeal in the form of electromagnetic waves. The phase and energy level of the electromagnetic waves both recover to the state prior to excitation. The image may be reconstructed after the echo signals transmitted by the nuclei are subjected to further processing (e.g., space encoding).
A magnetic resonance imaging system may operate with a number of various radio-frequency (RF) antennae (hereinafter, coils). A radio-frequency antenna is used to transmit and receive radio-frequency pulses so as to excite the atom nucleus to radiate magnetic resonance signals and/or to collect the induced magnetic resonance signals. The MRI system includes various coils, such as a body coil that covers the whole body area, a local coil that covers only a part of the body, or the like. The magnetic resonance system may have a large integrated coil (e.g., body coil) that is permanently fixed in a magnetic resonance scanner. The integrated coil may be arranged in a cylindrical manner around a patient sampling and collecting cavity (e.g., using a structure referred to as a nest configuration). In the patient sampling and collecting cavity, the patient is supported on a bed (e.g., a patient positioning table) during measurement. Since the coverage area of the body coil may be large, a higher transmitting power is used. The signal-to-noise ratio of an obtained image is relatively low, and the signal-to-noise ratio throughout the image is non-uniform. By contrast, the coverage area of a local coil may be small (e.g., the knee area covered by a knee coil, the head covered by a head coil, a wrist covered by a wrist coil, etc.). Thus, the local coil receives only limited radio-frequency signals within the radio-frequency excitation range. In order to distinguish the radio-frequency signals received by the local coil from the radio-frequency signals of the transmission stage, the radio-frequency signals received by the local coil are hereinafter referred to as magnetic resonance signals. The signal-to-noise ratio of an obtained image from a local coil may be high, and the signal-to-noise ratio throughout the image may be substantially uniform.
FIG. 1 shows a schematic diagram of a conventional coil control device of a magnetic resonance imaging system. In order to protect the security of the patient and the reliability of the coil itself, as shown in FIG. 1, the coil control device switches between a linear DC power supply (e.g., driven by a 15 V voltage VCC) and a negative voltage VSS (e.g., −32 V) for control. However, there are too many energy losses associated with conventional designs. A load at one side of the coil is a diode. A sink current ICS passing through the diode is several hundreds of milliamperes with regard to the coil and several amperes with regard to the body coil. The energy loss produced from a control circuit may be obtained according to the following formula,PDISS=ICS(VCC−VF),where YF is a forward voltage passing through the diode (e.g., about 0.7 V). In this example, the energy loss rate may be shown by the following formula:η=PDISS/PTTL*100%=ICS(VCC−VF)/(ICSVCC)*100%,where PTTL is the whole energy. In the example where VCC is 15 V and VF is 0.7 V, the energy loss rate η reaches 93.5%.
As shown above, in such a coil control circuit, most of the energy is wasted, thereby resulting in over-heating of the coil control circuit. As a result, a water-cooling system is used with the coil control circuit, thereby occupying space and increasing costs. In addition, the coil control circuit uses two power supplies that are positive and negative, thereby further increasing system complexity and costs.